X-ray systems are employed for imaging for diagnostic examination and for interventional procedures, (e.g., in cardiology, radiology and also surgery). FIG. 1 depicts such a monoplanar x-ray system presented by way of example, with a C-arm 2, held by a pedestal 1, in the form of a six-axis industrial or articulated-arm robot, to the ends of which an x-ray radiation source, (for example, an x-ray emitter 3 with x-ray tube and collimator, and an x-ray image detector 4 as image capturing unit), are attached.
By the articulated-arm robot, for example, from U.S. Pat. No. 7,500,784 B2, which may have six axes of rotation and thus six degrees of freedom, the C-arm 2 may be adjusted in any given spatial direction, for example, by the arm being rotated around a center of rotation between the x-ray emitter 3 and the x-ray image detector 4. The angiographic x-ray system 1 to 4 is able to be rotated around centers of rotation and axes of rotation in the C-arm plane of the x-ray image detector 4, e.g., around the center point of the x-ray image detector 4 and around the axes of rotation intersecting the center point of x-ray image detector 4.
The articulated-arm robot has a base frame that is mounted fixed to the floor, for example. Attached thereto is a carousel able to be rotated around a first axis of rotation. Attached to the carousel is a robotic motion link able to be pivoted around the second axis of rotation, to which a robot arm is fastened able to be rotated around a third axis of rotation. A robot hand able to be rotated around the fourth axis of rotation is attached to the robot arm. The robot hand has a fastening element for the C-arm 2, which is able to be pivoted around a fifth axis of rotation and rotated around a sixth axis of rotation running at right angles thereto.
The realization of the x-ray diagnostic device is not dependent on the industrial robot. Other conventional C-arm devices may also be used.
A patient 6 to be examined located on a table plate 5 of a patient support table as examination object is located in the beam path of the x-ray emitter 3. Connected to the x-ray diagnostic device is a system control unit 7 with a high-voltage generator for generating the tube voltage and an image system 8 that receives the image signals of the x-ray image detector 4 and processes them. The x-ray images may then be viewed on displays of a monitor stand 10 held by a ceiling-mounted, longitudinally-movable, pan and tilt and height-adjustable carrier system 9. Also provided in the system control unit 7 is a processing circuit 11, the function of which will be described below.
Instead of the x-ray system depicted by way of example in FIG. 1 with the stand 1 in the form of the six-axis industrial or articulated-arm robot, the angiographic x-ray system may also have a normal ceiling-mounted or floor-mounted holder for the C-arm 2.
Instead of the C-arm 2 depicted by way of example, the angiographic x-ray system may also have separate ceiling-mounted and/or floor-mounted holders for the x-ray emitter 3 and the x-ray image detector 4, which are electronically rigidly coupled for example.
Biplanar x-ray systems with two planes (e.g., two C-arms) may likewise be employed in interventional radiology.
The x-ray emitter 3 emits a ray bundle 12 emanating from a beam focus of its x-ray radiation source that strikes the x-ray image detector 4. If three-dimensional (3D) datasets are to be created in accordance with what is referred to as the DynaCT method, a method for rotation angiography, the rotatably supported C-arm 2 with x-ray emitter 3 and x-ray image detector 4 is rotated such that, as is depicted schematically from above onto the axis of rotation in FIG. 2, the x-ray emitter 3 depicted here graphically by its beam focus, as well as the x-ray image detector, 4 move around an object 13 to be examined located in the beam path of the x-ray emitter 3 on an orbital track 14. The arm may move around the orbital track 14 completely or partly to create a 3D data record.
The C-arm 2 with x-ray emitter 3 and x-ray image detector 4 may move in this case, in accordance with the DynaCT method, around at least one angular range of 180°, (e.g., 180° plus beam angle), and records a rapid succession of projection images from different projections. The images may only be reconstructed from a part range of this recorded data.
The object 13 to be examined may for example involve an animal or human body but also a phantom body.
The x-ray emitter 3 and the x-ray detector 4 move around the object 5 in each case so that the x-ray emitter 3 and the x-ray detector 4 are opposite one another on opposing sides of the object 13.
In normal radiography or fluoroscopy by such an x-ray diagnostic device, the medical two-dimensional (2D) data of the x-ray image detector 4 might be buffered in the image system 8 and subsequently reproduced on the monitor 9.
Flat-panel x-ray detectors may be used for full coverage in many areas of medical x-ray diagnostics and intervention, (e.g., in radiography, interventional radiology, or cardio angiography), but also therapy for imaging as part of the checking and radiation planning or mammography.
Current flat-panel x-ray detectors may be integrating detectors and are based primarily on scintillators, of which the light is converted in matrixes of photodiodes into electrical charge. These may be read out via active control elements row-by-row.
FIG. 3 depicts the basic structure of an embodiment of an indirect converting x-ray flat-panel detector 4 used, including a scintillator (CsI) and active read-out matrix (photodiode and (TFT) switching element) in a perspective cross-section. The core components of the x-ray image detector 4 include a solid-state pixel matrix, row drivers and amplifiers, as is described for example in “Flat detectors and their clinical applications” by Martin Spahn, Eur Radiol. (2005), Vol. 15, pages 1934 to 1947. The solid-state pixel matrix may have a layer with a scintillator 15, including cesium iodide (CsI) for example, which on irradiation by x-ray radiation, feeds visible photons into a pixel matrix 16 of amorphous silicon (aSi), that produce a visible x-ray image. Each of the plurality of pixels or image points of this pixel matrix 16 includes, as is depicted enlarged in FIG. 3, of a pixel element 17 with a photodiode 18 and a switch 19, which is connected to row lines 20 and column lines 21. The pixel matrix 16 is applied to a glass substrate 22.
All pixel elements of a row are addressed and read out at the same time by the activation electronics 23 in each case. In the simplest case, an image is read out progressively row-by-row. The signals are fed via readout electronics 24 to the processing circuit 11 disposed in the imaging system 8, for example, in which the signals are processed in parallel in a plurality of amplifiers, merged by multiplexers and converted in an analog/digital converter (A/D converter) into a digital output signal for further digital processing.
Depending on the beam quality, the quantum efficiency for a scintillator made of CsI with a layer thickness of, e.g., 600 μm is between around 50% and 80%. The local-frequency dependent DQE(f) (“detective quantum efficiency”) is limited upwards by this and for pixel sizes, (e.g., 150 μm to 200 μm) and for the local frequencies of 1 to 2 lp/mm of interest for the applications lies well below this figure. To make new clinical applications possible, (e.g., dual-energy, spectral imaging, material separation), but also to further increase the quantum efficiency, the potential of counting detectors for energy-discriminating counting detectors mainly based on direct-converting materials such as CdTe or CdZnTe (CZT) and contacted ASICs (application specific integrated circuits; e.g., implemented in CMOS technology) are increasingly being investigated.
The principal structure of such counting x-ray image detectors 4 is depicted in FIG. 4, which depicts a schematic diagram of a side view of a direct-converting, counting x-ray image detector 4 including a number of detector modules 25. Each detector module 25 includes a flat direct converter 26, (e.g., CdTe or CZT), which converts the x-ray radiation and separates the created charge carrier pairs via an electrical field that is created by a common top electrode 27 (e.g., cathode) and a pixel electrode 28 (e.g., anode). The charging creates a charge pulse in one of the pixel electrodes 28 of a (e.g., pixelated) ASIC 29 designed in pixel form, the height of which corresponds to the energy of the incident x-ray quantum and which, if lying above a defined threshold value, is registered as a counter event. The threshold value serves to distinguish an actual event from electronic noise or, e.g., also to suppress k fluorescence photons, in order to distinguish multiple counting. Disposed between the pixel electrodes 28 of the ASIC 29 in each case is pixel electronics 30, which provides at least one pulse shaper, comparator, counter, and switching electronics for the readout process, as will be to some extent explained in greater detail below. The ASIC 29, a corresponding section of the direct converter 26 and a coupling between direct converter 26 and ASIC 29 (for direct converting detectors, e.g., by contacts 31 known as bump bonds) each form the detector module 25 with a plurality of pixel elements 17. The ASIC 29 is disposed on a substrate 32 and connected, for example, by TSV connections 33 (e.g., through silicon via) to peripheral electronics 34. A detector module 25, if required, may also have one or more ASICs 29 and one or more parts sections of the direct converter 26. A high voltage connection 35 is routed to the common top electrode 27. The top electrodes 27 of the detector module 25 as well as if necessary the parts sections of the direct converter 26 are in contact via high voltage connections 36. The peripheral electronics 34 is connected to at least one tile, a detector module 25, or, in certain embodiments, to a number of tiles or detector modules 25.
FIG. 5 depicts a three-dimensional view of such a structure. For large-area x-ray image detectors 4 (e.g., 20×30 cm2), a number of detector modules 25 each of 2×2 cm2 surface area, for example, are connected together (e.g., 10×15 of such detector modules 25 would thus be provided) and linked via the common peripheral electronics 34. The connection between the ASIC disposed on the substrate 32 and the pixel electrodes 28 and the pixel electronics 30 and the peripheral electronics 34 are made by the TSV connections 33 (e.g., through silicon via).
FIG. 6 depicts the schematic layout of a counting pixel element 37. It may include the detector input, a preamplifier, a pulse shaper, a discriminator of which the threshold Vt may be configured, a counter, and an activation and readout unit. The electrical charge passes through the charge or signal input 38 as detector input. The electrical charge is collected in the pixel element 37 and is amplified there with the aid of a charge amplifier 39 and a feedback capacitor 40. In addition, at the output, the pulse shape may be configured in a shaper or filter. An event is counted when the output signal lies above an adjustable threshold. This is established via a discriminator 41. The threshold may, in principle, also be predetermined in a similar way as a fixed value by a threshold generator 42, but may be applied via a digital-to-analog converter (D/A converter, DAC) and is thus variably adjustable in a certain range. If the threshold is exceeded a digital counter 43 is incremented by one. Subsequently, the value may be read out via readout logic 44.
FIG. 7 depicts a corresponding schematic section of an ASIC module or array 45 of counting pixel elements 37, e.g., 100×100 pixel elements 37 each with an edge length of 180 μm, for example. Such an array 45 is realized with the aid of the ASIC 29. In this example, for a detector module 25, this would produce a size of 1.8×1.8 cm2. For larger x-ray image detectors 4 (e.g., 20×30 cm2), a number of detector modules 25 are combined (e.g., 11×17 detector modules 25 would produce approximately this surface area) and connected via a common peripheral control and readout unit 46 that implements a readout 47.
In the case of counting and energy-discriminating x-ray detectors, two, three, or more thresholds are introduced and the level of the charge pulse, corresponding to the predefined thresholds (e.g., discriminator thresholds), are arranged in one or more of the digital counters. The x-ray quanta counted in a specific energy range are able to be obtained by differentiating the counter contents of two corresponding counters. The discriminators are able to be set with the aid of the threshold generator for the entire detector module or pixel-by-pixel within given limits or ranges. The counter contents of the pixel elements are read out one after the other via a corresponding readout facility. This readout process requires a certain time during which the count may not continue without errors.
FIG. 8 depicts a further example for the central function elements of a counting pixel element 37 of a digital x-ray image detector 4 with energy-discriminating pixel design. In this example, the pixel element 37 has several, (e.g. four), discriminators 41.1, 41.2, 41.3 and 41.4 each with a threshold generator 42.1, 42.2, 42.3 and 42.4 with different thresholds, to which digital counters 43.1, 43.2, 43.3, 43.4 are connected. Such a structure with the use of different discriminator thresholds provides an energy-selective and energy-discriminating imaging to be performed.
Different effects may now lead to an absorbed x-ray quantum not only depositing its energy in one pixel element 37 but, through processes such as charge sharing or fluorescence photons (k-fluorescence), to a part of the energy being deposited in the neighboring pixel elements 37. This is depicted schematically in FIG. 9, on the basis of which the circumstances during x-ray imaging will now be described in greater detail.
A primary x-ray quantum 50 falling on the x-ray image detector 4 creates a primary event 51 in the direct converter 26, which leads to a signal portion 52 at the primary location in a first pixel element 53. Through charge sharing, further signal portions 55 arise in the neighboring pixel elements 54. A fluorescence photon 56 may also be triggered by the primary event 51, which leads to a secondary event 57 in a direct converter 26 assigned to a neighboring pixel element 54, which results in a signal portion 58 through fluorescence photons 56 in the neighboring pixel element 54. This may lead to miscounting, for example to multiple counting or to no counting at all, if the respective deposited energies lie below the thresholds. The energy may also be incorrectly assigned in the case of energy-discriminating counting detectors.